Biomaterials are permanently or temporarily used to repair or replace missing or diseased parts of the human body (e.g. joint replacements, contact lenses, heart valves, vascular prostheses, dental implants, etc). Ultimately, almost every human in technologically advanced societies will host a biomaterial.
Ceramics are extremely popular in medical and dental applications because of their strength, chemical inertness, wear resistance, and esthetics. However, the full potential of ceramics in biomedical applications has not been realized, with biointergration and/or strength being the major concerns. The strong ceramics such as zirconia and alumina are bioinert and possess high elastic modulus. The bioinert property prevents the formation of chemical bonds with the surrounding tissues. The high elastic modulus results in stress shielding of the hard tissue, leading to local bone resorption. The bioactive ceramics (e.g. calcium phosphate ceramics, bioglasses) are able to form strong chemical bonds with adjacent tissues and exhibit comparable modulus to hard tissues (bone, enamel, or dentin). Unfortunately, bioactive ceramics are relatively weak and thus structurally unstable.
The initial applications of ceramics in medicine are based on their chemical inertness and wear resistance. (Hench et al., Science 2002, 295(5557):1014-7) However, the inert surface fails to form biochemical bonds with the surrounding tissues, often leading to implant loosening. In addition, the high elastic modulus of the strong, bioinert ceramics (often an order of magnitude higher than that of hard tissues) results in stress shielding of the surrounding bone, leading to local bone resorption. (Liu et al., 2004, 47(3-4):49-121) The development of the bioactive ceramics in the late 1960s and early 1970s (e.g. calcium phosphate ceramics, silica-based bioactive glasses and glass ceramics) allow their use where bonding to hard tissues is needed. (Hench et al., Journal of Biomedical Materials Research 1971; 5:117-41; LeGeros, Monogr Oral Sci 1991, 15:1-201).
In addition, the bioactive ceramics have a relatively low modulus (on the same order of hard tissues) and therefore do not have a significant stress shielding problem. However, the bioactive ceramics exhibit relatively low strength and fracture toughness, and thus they are not suitable for load bearing applications (e.g. dental and orthopedic implants).
Public demands in biological compatibility and esthetics have driven the increasing popularity of ceramics as the materials of choice for dental and orthopedic implant devices. Dental and orthopedic implants bear load during function, requiring strong ceramics. However, strong ceramics such as zirconia and alumina are bioinert, failing to form strong bonds with vital tissues. In addition, strong alumina and zirconia ceramics possess high modulus of elasticity, resulting in local bone resorption. Both the bioinertness and high modulus can lead implant loosening and failure. On the other hand, the low modulus, bioactive ceramics are relatively weak and thus structurally unstable.
The current approaches to this problem are to either make a composite consisting of both inert and active phases, e.g., mixed hydroxyapatite (HA) and zirconia or alumina, or to coat the surface of a strong and inert ceramic with a bioactive but weak layer (e.g., plasma-sprayed HA coated implants, bioglass coated implants). (Inuzuka et al., Solid State Ionics 2004, 172(1-4):509-513; Moon et al. Eco-Materials Processing & Design Vi 2005, 486-487:101-104; Sun et al., J Biomed Mater Res 2001, 58(5):570-92). Compared to their parental materials, the composites exhibit a moderate bioactivity and strength. On the other hand, the bioactive ceramic coatings on the biologically inert ceramic are likely to undergo delamination and fracture due to the coating/substrate bonding issue, coefficient of thermal expansion (CTE) mismatch, and the abrupt change in physical and mechanical properties at the coating/substrate interface.
Coating of inert materials with bioactive glasses dates to the mid 1970s, with a primary effort in modifying the surface of metallic orthopedic stems and dental implants for better bone integration. (Hench et al., An Introduction to Bioceramics. 1993). The strong, tough metallic substrate provides structural support to the weak, brittle bioglass coating, while the bioglass coating protects the surrounding tissues from corrosion products of the metal core which may induce systemic effects. (Black J., J Biomed Mater Res 1976, 10(4):503-9). The problem of obtaining a bioactive glass coating with high mechanical integrity is the chemical reactivity of this type of glass. (Hench et al., An Introduction to Bioceramics. 1993). Silicate-based bioactive glasses contain less than 60 mol % SiO2, and thus exhibit a random two-dimensional sheet-like network structure with many open pathways for ion transport. The open network structure facilitates the rapid formation of a calcium hydrocarbonate apatite layer at the glass surface, which provides binding sites to bone or soft tissue. However, it is also this open network structure that provides pathways for other cations, such as Fe, Cr, Ni, Co, Mo, Ti, or Ta, to migrate from the metal substrate to the glass surface. The presence of these cations at the glass surface inhibit or eliminate the bioactivity of the glass by preventing formation of the HA layer.
Since oxide ceramics are chemically more stable and biologically more compatible than metals and alloys, attempts have been made to coat structural oxide ceramics, such as alumina and zirconia, with bioglasses. (Greenspan et al., J Biomed Mater Res 1976, 10(4):503-509; Ferraris et al. Biomaterials 2000, 21 (8):765-73). The major problem with these systems is a large difference in CTE: the silicate-based bioglasses have a CTE ranging between 12−16×10−6, while alumina and zirconia have CTE values of 8×10−6 and 10.5×10−6, respectively. The high CTE value of bioglass places the coating in tension, which further weakens the bioglass coating. One solution to the above mentioned problems is to utilize two layers of glass coatings. One laboratory developed two layer coatings for Co—Cr—Mo alloy and alumina substrates, and another laboratory exploited multilayer coatings for titanium alloy. (Lacefield et al., Biomaterials 1986, 7(2): 104-8; Gomez-Vega et al., Advanced Materials 2000, 12(12):894-898)
For example, recent reports of using silica-based bioactive coating on titanium alloy (Ti6Al4V) implants show improved bioactivity compared to the original Ti6Al4V surface. (Gomez-Vega et al., Journal of Dental Research 1998, 77:108-108). However, adhesion of the silica-based bioglass coatings on Ti6Al4V surfaces relies on the silica content. A high silica content forms a better bond with the Ti6Al4V surface, but also sacrifices the bioactivity of the glass coating. Therefore, multiple coating layers are required to grade the silica content in order to retain the necessary surface bioactivity and a sufficient interfacial bond with the Ti6Al4V substrate. Even then, cracks may be observed in the outer bioactive glass layer due to a large CTE mismatch between the glass coatings and Ti6Al4V substrate. One in vitro study revealed that the Ca, P-rich surface layer separated from the underlying glass due to the degradation of the silica network in bioglass. (Foppiano et al., Acta Biomaterialia 2006, 2(2):133-142). In addition, the silica based bioglass upon dissolution produce a basic environment (pH approximately 9-11), hindering tissue integration. CPG is chemically more stable (due to the absence of silica) and biologically more active (chemically more close to hard tissues) than silica-based bioglass.
Despite significant improvements that have been made to bond bioglass coating to metals and ceramics, widespread application of bioglass coated dental and orthopecdic implants has failed because of the fracture of glass coatings, making them poor candidates for load bearing applications.
It would be useful to create a functionally graded CPG/Y-TZP system with a low modulus, bioactive surface and yet a flexural strength similar to, or even greater than Y-TZP for dental and orthopedic implants. The osteoconductive CPG coating promotes a rapid osteointegration and prevents micromotion at the implant/tissue interface, while the graded CPG/Y-TZP structure retains excellent contact and flexural damage resistance. In addition, the residual outer surface CPG layer acts as an encapsulation layer, preventing hydrothermal degradation of Y-TZP interior, and can be further transformed to a carbonate apatite (CHA) layer by immersing in calcifying solution or simulated body fluid (SBF) with an electrolyte composition similar to that of serum since in all bioactive materials (e.g., calcium phosphates, bioactive glass, calcium sulfates, etc), the newly formed bone is directly attached to a CHA layer. Knowledge generated from this investigation can readily be extended to development of next-generation, strong ceramic scaffolds for medical applications, foreshadowing an array of engineering applications.